System and method for providing biomechanically suitable running gait in powered lower limb devices

ABSTRACT

Systems and methods for a running controller for a lower limb device including at least a powered knee joint are provided. The method includes collecting real-time sensor information for the lower limb device and configuring the lower limb device to a first state in a finite state model for an activity mode including the running mode. The method further includes, based on the sensor information, transitioning the lower limb device from a current state to a subsequent state in the finite state model for the detected mode when a pre-defined criteria for transitioning to the subsequent state is met, and repeating the transitioning until the activity mode changes. In the system and method, the finite state model includes at least one stance state and at least one swing state, where the at least one stance state includes at least one absorption state and at least one propulsion state.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application claims priority to and the benefit of U.S. ProvisionalApplication No. 61/610,864, filed Mar. 14, 2012 and entitled “CONTROLMETHODOLOGY FOR BIOMECHANICALLY NORMAL RUNNING WITH A POWEREDPROSTHESIS”, the contents of which are herein incorporated by referencein their entirety.

FIELD OF THE INVENTION

The present invention relates to apparatus and methods for controllinglower limb devices, including prostheses or orthoses, in order toprovide a biomechanically suitable running gait in such devices.

BACKGROUND

In 2005, approximately 623,000 cases of lower limb amputation existed inthe United States, with the total number of cases of limb loss expectedto increase by approximately 40% by the year 2020. Lower-limb prosthesesexist in large part to improve the mobility of the user, particularlyconcerning activities involved in daily living. Dedicated sportsprostheses also exist for activities such as running, field events,cycling, swimming, golf, etc., which are used for competition andrecreation alike. Many of these devices have been proven quiteeffective, some having even been accused of providing an unfairadvantage to the user. However, should the need arise during the courseof normal daily activity for a lower-limb amputee to run—perhaps toquickly dodge an oncoming vehicle or to catch a bus—the individual'sdaily use prosthesis would be called upon to meet that need.

The majority of prostheses currently available to lower limb amputeesare energetically passive. Passive prostheses are unable to reproducethe biomechanics of healthy running in part because these biomechanicsrequire significant net positive power at both the knee and anklejoints. In recent years, powered lower limb prostheses, which are ableto produce net positive power at the knee or ankle joints, have startedto emerge. However, none of these devices incorporate both a poweredknee and ankle (defined as able to produce biomechanically significantnet positive power at each joint over a stride). Moreover, none of thesedevices have demonstrated restoration of healthy gait characteristicsfor running in transfemoral amputees. Relative to walking,biomechanically healthy running is characterized by a substantiallygreater degree of stance knee flexion and a correspondingly greaterdegree of ankle dorsiflexion, also in the stance phase. Further, thestance phase of running constitutes less than 50% of the stride cycle,while the stance phase of walking constitutes greater than 50%. As such,a walking gait is typically characterized by a double support phase,while a running gait is typically characterized by a double float (orflight) phase, i.e., a phase in which both feet are off the ground. Inorder to provide the latter, each leg must generate an amount ofvertical propulsive energy greater than or equal to the amount absorbedduring each stance phase of gait. Since the foot/ground collision willalways realistically involve some energy loss, each leg must in factgenerate an amount of propulsive energy strictly greater than the amountabsorbed. In order to do so, the joints of a running leg capable ofsustaining a running gait must be powered (i.e., they must be capable ofgenerating more power than they absorb).

SUMMARY

Embodiments of the invention concern systems and methods for controllinglower limb devices.

In a first embodiment of the invention, a method of operating a lowerlimb device comprising at least a powered knee joint is provided. Themethod includes collecting real-time sensor information for the lowerlimb device and configuring the lower limb device to a first state in afinite state model for an activity mode comprising the running mode. Themethod also includes, based on the sensor information, transitioning thelower limb device from a current state to a subsequent state in thefinite state model for the detected mode when a pre-defined criteria fortransitioning to the subsequent state is met. The method furtherincludes repeating the transitioning until the activity mode changes. Inthe method the finite state model comprises at least one stance stateand at least one swing state, and wherein the at least one stance statecomprises at least one absorption state and at least one propulsionstate.

In the method, the transitioning can include causing the powered kneejoint to dissipate an first amount of power during the at least oneabsorption state, and causing the powered knee joint to generate asecond amount of power during at least one propulsion state, where thesecond amount of power is at least equal to the first amount of power.The transitioning can further include, when the at least one poweredjoint further comprises a powered ankle joint, causing the powered kneejoint and the powered ankle joint to simultaneously dissipate powersubstantially throughout the at least one absorption state, and causingthe powered knee joint and the powered ankle joint to simultaneouslygenerate power substantially throughout the at least one propulsionstate.

In the method, the transitioning can also include causing the at leastone powered joint to emulate a passive impedance during the at least oneabsorption state. The passive impedance can be at least one of astiffness component or a damping component.

In the method, the pre-defined criteria associated with a transitionbetween the at least one absorption state and the at least onepropulsion state can be associated with at least one of a motion in atleast one powered joint of the lower limb device, a joint angularvelocity for at least one powered joint of the lower limb device, andload on the lower limb device.

The method can further include, prior to the configuring, selecting theactivity mode for the lower limb device based on the real-time sensorinformation, where a transition between a walking mode and the runningmode is based on a measurement of at least one of a load or accelerationat foot strike, a stance time, a swing time, or a stride time.

In a second embodiment of the invention, there is provided acomputer-readable medium having stored thereon a plurality ofinstructions for causing a controller device for a powered lower limbdevice to perform any of the methods of the first embodiment.

In a third embodiment of the invention, there is provided a system forcontrolling a lower limb device comprising at least a powered kneejoint. The system includes at least one sensor for collecting real-timesensor information for the lower limb device and at least one processorcommunicatively coupled to the at least one sensor and to the at leastpowered knee joint. The system also includes a computer-readable medium,having stored thereon instructions for causing the processor to performvarious steps. The steps include generating control signals for at leastthe powered knee joint to transition the lower limb device to a firststate in a finite state model for an activity mode comprising a runningmode and generating additional control signals for at least the poweredknee joint to transition the lower limb device from a current state to asubsequent state in the finite state model when a pre-defined criteriafor transitioning to the subsequent state is met based on the real-timesensor information. The steps also include repeating the generating ofthe additional control signals transitioning until the activity modechanges. In the system, the finite state model comprises at least onestance state and at least one swing state, and wherein the at least onestance state comprises at least one absorption state and at least onepropulsion state.

In the system, the instructions can cause the processor to generate thecontrol signals for the at least one absorption state to cause thepowered knee joint to dissipate an first amount of power during the atleast one absorption state and to generate the control signals for theat least one propulsion state to cause the powered knee joint togenerate a second amount of power during the at least one propulsionstate, and where the second amount of power is at least equal to thefirst amount of power. Further, where the at least one powered jointfurther comprises a powered ankle joint, the instructions can cause theprocessor to generate the control signals for the at least oneabsorption state to cause the powered knee joint and the powered anklejoint to simultaneously dissipate power substantially throughout the atleast one absorption state and to generate the control signals for theat least one propulsion state to cause the powered knee joint and thepowered ankle joint to simultaneously generate power substantiallythroughout the at least one propulsion state.

The instructions can also cause the processor to cause the at least onepowered joint to emulate a passive impedance during the at least oneabsorption state. The passive impedance can be at least one of astiffness component or a damping component.

In the system, the pre-defined criteria associated with a transitionbetween the at least one absorption state and the at least onepropulsion state is associated with at least one of a motion in at leastone powered joint of the lower limb device, a joint angular velocity forat least one powered joint of the lower limb device, and load on thelower limb device.

Further, the computer-readable medium further comprises instructions forcausing the processor to perform the step of, prior to the generating ofthe control signals, selecting the running mode for the lower limbdevice based on the real-time sensor information, wherein a transitionbetween a walking mode and the running mode is based on a measurement ofat least one of a load or acceleration at foot strike, a stance time, aswing time, or a stride time.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1A is a view of a powered knee and ankle prosthesis, according toan embodiment of the invention;

FIG. 1B is an exploded view of the powered knee and ankle prosthesisshown in FIG. 1A, according to an embodiment of the invention;

FIG. 2 is an exploded view of knee motor unit, according to anembodiment of the invention;

FIG. 3 is an exploded view of ankle motor unit, according to anembodiment of the invention;

FIG. 4 is an exploded view of knee joint, according to an embodiment ofthe invention;

FIG. 5 is an exploded view of ankle joint, according to an embodiment ofthe invention;

FIGS. 6A and B are views of a foot having toe and heel force sensingelements, according to an embodiment of the invention;

FIG. 7 shows the joint angle and torque convention used herein. Positivetorque is defined in the direction of increasing angle;

FIG. 8 shows the subdivision of normal walking into four internal phasesshowing the knee and ankle angles during the phases, according to anembodiment of the invention;

FIG. 9 shows a finite-state model of normal walking, according to anembodiment of the invention. Each box represents a different internalphase and the transition conditions between the internal phases arespecified;

FIG. 10 shows piecewise fitting of knee and ankle torques during normalspeed level walk scaled for a 75 kg adult to a non-linear spring-damperimpedance model;

FIG. 11 is a diagram for an active/passive decomposition based controlof the powered knee and ankle prosthesis, according to an embodiment ofthe invention;

FIG. 12 is a diagram for a general form of active-passive decompositioncontrol including intent recognition that provides supervisorymodulation, according to an embodiment of the invention;

FIG. 13A is a side view of powered knee and ankle prosthesis, accordingto another embodiment of the invention;

FIG. 13B is a front view of powered knee and ankle prosthesis of FIG.13A.

FIGS. 14A and 14B show perspective and bottom views of an exemplarysagittal moment load cell suitable for use in the various embodiments ofthe invention;

FIG. 15 is a block diagram of an exemplary embedded microcontroller inaccordance with an embodiment of the invention;

FIG. 16 is a control state chart for the three activity modescorresponding to walking, standing, and sitting, and for the internalphases and their corresponding transitions within each activity mode;

FIG. 17 shows knee angle modulated knee stiffness during pre-stand(solid line) and pre-sit (dashed line) phases;

FIG. 18 is a plot of axial actuation unit force versus ankle angle;

FIG. 19 shows a normal speed walking phase portrait of the knee jointand four stride segments;

FIG. 20 shows the selection of indexing data samples during a firstsegment of a walking stride;

FIGS. 21A, 21B, and 21C are the output of the decomposition for Segment1 showing, respectively, the spring constants, the dashpot constants,and the active and passive knee torques;

FIG. 22 is a state chart for governing the discrete dynamics of anactive-passive decomposition controller in accordance with an embodimentof the invention;

FIG. 23 is a state chart for governing the discrete dynamics of thecadence estimator in accordance with an embodiment of the invention;

FIG. 24 is a schematic diagram of accelerometer measurements for slopeestimation in accordance with an embodiment of the invention;

FIG. 25 is a state chart for slope estimation in a controller inaccordance with an embodiment of the invention;

FIGS. 26A and 26B show front and back views of a friction/cable drivemotor in accordance with an embodiment of the invention;

FIG. 27 shows an exemplary embodiment of a belt drive transmission inaccordance with an embodiment of the invention;

FIGS. 28A and 28B show side views of first and second positions,respectively, achievable for an exemplary embodiment of a chain drivetransmission including an eccentric mount in accordance with anembodiment of the invention;

FIG. 29 illustrates schematically the components for the adjustablebearing mounts in FIGS. 28A and 28B;

FIG. 30 illustrates an exemplary configuration of a powered legprosthesis in accordance with the embodiments shown in FIGS. 27-29;

FIG. 31 shows the (body-mass-normalized) power characteristics of theknee and ankle joints during the stance phase of running for healthysubjects;

FIG. 32 shows the (body-mass-normalized) power characteristics of theknee and ankle during the stance phase of walking for healthy subjects;

FIG. 33 shows an exemplary running controller for a powered prosthesiswith a knee and ankle joint;

FIG. 34 shows a specific implementation for the exemplary runningcontroller of FIG. 33;

FIG. 35 depicts six key elements of a stride captured from a video takenduring one trial;

FIG. 36 depicts the mode transitions (percent of stride)±one standarddeviation as recorded during the running controller evaluations; and

FIG. 37 compares thirteen strides of the amputee subject running on thepowered prosthesis to the sagittal plane knee and ankle joint angles ofhealthy subjects.

DETAILED DESCRIPTION

The present invention is described with reference to the attachedfigures, wherein like reference numerals are used throughout the figuresto designate similar or equivalent elements. The figures are not drawnto scale and they are provided merely to illustrate the instantinvention. Several aspects of the invention are described below withreference to example applications for illustration. It should beunderstood that numerous specific details, relationships, and methodsare set forth to provide a full understanding of the invention. Onehaving ordinary skill in the relevant art, however, will readilyrecognize that the invention can be practiced without one or more of thespecific details or with other methods. In other instances, well-knownstructures or operations are not shown in detail to avoid obscuring theinvention. The present invention is not limited by the illustratedordering of acts or events, as some acts may occur in different ordersand/or concurrently with other acts or events. Furthermore, not allillustrated acts or events are required to implement a methodology inaccordance with the present invention.

In view of the limitations of existing lower limb prostheses and otherlower limb devices, the various embodiments are directed to a newcontrol algorithm (i.e., a running controller) that enables abiomechanically appropriate or suitable running gait for lower limbdevices. When such a running controller is implemented in a poweredprosthesis comprising at least a powered knee joint and optionally apowered ankle joint, the incorporated running controller can enableamputees with such a prosthesis to have a running gait closelyrepresentative of biomechanically healthy running, including theappropriate aforementioned joint kinematics and the double float phaseof gait.

Although the various embodiments will be described primarily withrespect to the incorporation of the running controller into a controlsystem for a lower limb prosthesis, the various embodiments are notlimited in this regard. Rather, the running controller can also beutilized in any type of lower limb devices, including, but not limitedto, powered orthotic devices or other lower limb assistive devices.Further, although the various embodiments will be described with respectto a lower limb prosthesis including a powered knee joint and a poweredankle joint (e.g., for transfemoral amputees), the various embodimentsare not limited in this regard. Rather, the running controller can alsobe used to improve running gait for lower limb prostheses for amputeeswith an intact knee joint (i.e., for transtibial amputees).

Prior to discussing the running, the disclosure first turns to FIGS.1A-30 where there are described various configurations for powered legand ankle prostheses, including a controller, which can be modified toinclude a running controller in accordance with the various embodimentsof the invention.

Exemplary Prosthesis Configurations

A first design for a prosthesis for use in the various embodiments ofthe invention is shown in FIG. 1A through FIG. 6B. The prosthesis 100comprises a prosthetic lower leg 101. Lower leg 101 can be coupled to apowered knee joint comprising a knee motor unit 105 coupled to a kneejoint 110, and a powered ankle joint comprising an ankle motor 115coupled to an ankle joint 120. A sagittal plane moment sensor 125 can belocated between the prosthesis and the user to measure the moment, andin one embodiment is located immediately below the socket interface. Inthe embodiment shown, sensor 125 measures the sagittal plane moment,while separate sensors described below measure the ball of foot forceand heel force with respect to the ground or other object the foot ispressed against. A load sensor 135 can be positioned at the ball of thefoot, and a load sensor 140 can be positioned at the heel of the foot.However, in another embodiment (not shown) sensor 125 can measure thesagittal plane moment, the frontal plane moment and the axial force,such as provided by the three-axis socket load cell. This alternateembodiment can eliminate the need for sensor 135 and sensor 140.

Load sensors 141 and 142 are in series with each motor unit 105 and 115,respectively for motor unit force control. Position sensors 151 and 152are provided at each joint 110 and 120 as shown in FIGS. 4 and 5respectively. Position sensors 151 measure joint angle (θ as used below)and can be embodied as potentiometers. The computer/process controller,and power source (e.g. a battery such as a Li ion battery, andelectrical connections in the case of an electrical power source are notshown to avoid obscuring aspects of the invention. Non-electrical powersources may also be used, such as pneumatic power, or non-batteryelectrical sources, such as hydrogen-based fuel cells.

Prosthesis 100 is shown in an exploded view in FIG. 1B. Joints 110 and120 are more clearly shown as compared to FIG. 1A.

FIG. 2 is an exploded view of knee motor unit 105, according to anembodiment of the invention. Load sensor 141 is shown as a load cell(e.g. strain gauge). Load sensor 141 measures force and moments. Themotor unit 105 comprises a motor-driven ball screw assembly which drivesthe knee joint through a slider-crank linkage comprising screw 145.Other motor drive assemblies may also generally be used.

FIG. 3 is an exploded view of ankle motor unit 115, according to anembodiment of the invention. Load sensor 142 is generally analogous toload sensor 141. The motor unit 115 comprises a motor-driven ball screwassembly which drives the ankle joint through a slider-crank linkagecomprising screw 145. The ankle motor 115 includes a spring 147positioned to provide power in parallel (thus being additive) with powerprovided by the motor unit 115. Spring 147 biases the motor unit's forceoutput toward ankle plantarflexion, and supplements the power outputprovided by motor unit 115 during ankle push off.

FIG. 4 is an exploded view of knee joint 110, according to an embodimentof the invention. As described above, knee joint 110 includes positionsensor 151 that can be embodied as a potentiometer for anglemeasurements of the knee joint 110.

FIG. 5 is an exploded view of ankle joint 120, according to anembodiment of the invention. As described above, ankle joint 120includes position sensor 152 that can be embodied as a potentiometer forangle measurements of the ankle joint 120.

FIG. 6A is a view of a foot 170 having ball of foot sensors 135,according to an embodiment of the invention. Sensors 135 are provided tomeasure the ground reaction forces near the ball of the foot, such aswhen the foot strikes the ground. FIG. 6B is a view of a foot 170 havingball of foot sensors 135 and heel sensors 140, according to anembodiment of the invention. Sensors 140 are provided to measure theground reaction forces on the heel of the foot when the foot 170 strikesthe ground. Sensors 135 and 140 can be embodied as strain based sensors.

Unlike existing passive prostheses, the introduction of power into aprosthesis according to embodiments of the invention provides theability for the device to also act, rather than simply react. As such,the development of a suitable controller and control methodology thatprovides for stable and reliable interaction between the user andprosthesis is provided herein. Control according to embodiments of theinvention has been found to enable the user to interact with theprosthesis by leveraging its dynamics in a manner similar to normalgait, and also generates more stable and more predictable behavior.

Thus, rather than gather user intent from the joint angle measurementsfrom the contralateral unaffected leg, embodiments of the inventioninfer commands from the user via the (ipsilateral) forces and moments ofinteraction between the user and prosthesis. Specifically, the userinteracts with the prosthesis by imparting forces and moments from theresidual limb to the prosthesis, all of which can be measured viasuitable sensor(s), such as sensors 125, 140 and 141 described abovewhich measures moments/forces. These forces and moments serve not onlyas a means of physical interaction, but also serve as an implicitcommunication channel between the user and device, with the user'sintent encoded in the measurements. Inferring the user's intent from themeasured forces and moments of interaction according to embodiments ofthe invention provides several advantages relative to the known echoapproach.

In one embodiment of the invention the torque required at each jointduring a single stride (i.e. a single period of gait) can be piecewiserepresented by a series of passive impedance functions. A regressionanalysis of gait data indicates that joint torques can be characterizedby functions of joint angle (θ) and angular velocity by an impedancemodel, such as the following exemplary passive impedance function shownin equation 1 below:τ=k ₁(θ−θ_(e))+b*{dot over (θ)}  (1)where k₁, b, and the equilibrium joint angle θ_(e) are all constantsthat are generally generated empirically, and are constants for a givenjoint during a given internal phase (e.g. knee, internal phase 3). k₁characterizes the linear stiffness, b is the linear damping coefficient,θ is the measured joint angle which can characterize the state of theprosthesis, θ_(e) is the equilibrium angle, {dot over (θ)} is theangular velocity of the joint, and τ is the joint torque. Given theseconstants, together with instantaneous sensor measurements for θ and{dot over (θ)}, the torque (τ) at the joints (knee and ankle) can bedetermined.

Positive directions of the angle (θ) and torque (τ) as used herein aredefined as shown in FIG. 7. If the coefficients b and k₁ are constrainedto be positive, then the joints will each exponentially converge to astable equilibrium at θ=θ_(e) and {dot over (θ)}=0 within each internalphase. That is, within any given internal phase, the actuators areenergetically passive (i.e. the joint will come to rest at a localequilibrium). While the unactuated prosthesis can be energeticallypassive, the behavior of one joint (knee or ankle) or the combinedbehavior of the knee and ankle joints, can be likewise passive, and thuswill generally respond in a predictable manner.

Responsive to direct input from the user (e.g. a heel strike) to triggera change in internal phase, power (torque) can be delivered from thepower source (e.g. battery) to the prosthesis in the proper magnitude toprovide the desired movement. Since the switching can be triggered bydirect input from the user related to the current internal phase, theuser maintains direct influence over the power applied to theprosthesis. If the user does not trigger the next internal phase (i.e.remains stationary) no net energy is delivered. That is, the prosthesiswill generally cease to receive power from the power source for movingthe joint, and will instead, due to the damped response, soon come torest at the local equilibrium identified with the present internalphase.

As described above, the decomposition of joint behavior into passivesegments requires the division of the gait cycle into a plurality ofinternal phases or “finite states” characterized by an impedancefunction and a set of constants for the impedance function, as dictatedby their functions and the character of the piecewise segments of theimpedance functions described above. The switching rules betweeninternal phases should generally be well defined and measurable, and thenumber of phases should be sufficient to provide a substantiallyaccurate representation of normal joint function. In one embodiment ofthe invention, the swing and stance phase of gait can constitute aminimal set of internal phases.

Based on least-squares regression fitting of Equation 1 to empiricalgait data, the present Inventors determined that such fits were improvedsignificantly by further dividing the two modes of swing and stance eachinto two additional internal phases to realize four phases, as shown inFIG. 8. A fifth internal phase can also be added, as illustrated in FIG.16. The angle (θ) of the prosthetic knee (above) and ankle joint (below)can be provided during each internal phase as a function of the % of thestride. Angle values shown can be used as threshold values to triggerphase changes as described below relative to FIG. 9. As clear to onehaving ordinary skill in the art, the number of phases can be other thantwo or four.

FIG. 9 shows exemplary switching rules between internal phases forwalking FIG. 16 shows another set exemplary switching rules, forwalking, standing, and sitting activity modes. As described above, ifthe user does not initiate actions that trigger the next phase (e.g.based on the switching rules), the prosthesis will cease to receivepower and will come to rest at the local equilibrium identified with thepresent phase. For example, switching can be based on the ankle angle>athreshold value (Mode 1 to Mode 2), or ankle torque<threshold) (Mode 2to Mode 3), the angle or torque measurements provided by on boardsensors as described above.

Phase 1 shown in FIG. 8 begins with a heel strike by the user (which canbe sensed by the heel force sensor), upon which the knee immediatelybegins to flex so as to provide impact absorption and begin loading,while the ankle simultaneously plantarflexes to reach a flat foot state.Both knee and ankle joints have relatively high stiffness (and can beaccounted for by k1 in equation 1) during this phase to prevent bucklingand allow for appropriate stance knee flexion, because phase 1 comprisesmost of the weight bearing functionality. Phase 2 is the push-off phaseand begins as the ankle dorsiflexes beyond a given angle (i.e. user'scenter of mass lies forward of stance foot). The knee stiffnessdecreases in this mode to allow knee flexion while the ankle provides aplantarflexive torque for push-off. Phase 3 begins as the foot leavesthe ground as detected by the ankle torque load cell and lasts until theknee reaches maximum flexion. Mode 4 is active during the extension ofthe knee joint (i.e. as the lower leg swings forward), which begins asthe knee velocity becomes negative and ends at heel strike (e.g. asdetermined by the heel force sensor).

In both of the swing phases (Phases 3 and 4), the ankle torque can besmall and can be represented in the controller as a (relatively) weakspring regulated to a neutral position. The knee can be primarilytreated as a damper in both swing phases.

Impedance modeling of joint torques was preliminarily validated byutilizing the gait data from a healthy 75 kg subject, as derived frombody-mass normalized data. Incorporating the four internal phasesdescribed above, along with the motion and torque data for each joint, aconstrained least-squares optimization was conducted to generate a setof parameters k₁, b and θ_(e) for each phase for each joint for use inEquation 1. The resulting parameter set can be fit to joint torques andis shown graphically in FIG. 10. FIG. 10 shows piecewise fitting of kneeand ankle torques during normal speed level walk scaled for a 75 kgadult to a non-linear spring-damper impedance model. The numbers shownin each phase represent the mean ratio of the stiffness forces todamping forces predicted by the fit. The vertical lines represent thesegmentation of a gait stride into four distinct phases. The fit shownin FIG. 10 clearly indicates that normal joint function can berepresented by the use of piecewise passive functions.

Controllers according to embodiments of the invention generally comprisean underlying gait controller (intra-modal controller). An optionalsupervisory gait controller (also called intent recognizer) can also beprovided. Both controllers generally utilize measured information. Thisinformation generally comprises user and ground interaction forces (F)and moments/torques (τ), joint angles and angular velocities fromon-board sensors, and can be used to extract real-time input from theuser. The gait control component utilizes the sensed instantaneousnature of the user input (i.e., moments and forces) to control thebehavior of the leg within a given activity mode, such as standing,walking, or stair climbing.

Two exemplary approaches to intra-modal impedance generation aredescribed below. The first approach is shown in FIG. 11 and represents ageneral form of active-passive decomposition-based intra-mode control.The second embodiment shown in FIG. 12 includes the control structureshown in FIG. 11 but adds a supervisory intent recognizing controller tomodulate the intra-modal control based on inputs from an intentrecognition module. As shown in FIGS. 11 and 12, F_(s) is the force theuser of the prosthesis is applying, such as a heel force in the case ofa heel strike, τ represents joint torque, and θ represent joint angles.τ_(a) represents the active component of joint torque which is roughlyproportional to the input force, and τ_(p) represents the passivecomponent of torque. The active joint torque τ_(a) is thus the totaljoint torque τ minus the passive joint torque, τp. Derivatives are shownusing the dot convention, with one dot being the first derivative (e.g.,{dot over (θ)} being angular velocity) and two dots representing thesecond derivative.

In the embodiment of the intra-modal controller shown in FIG. 11, thebehavior of the prosthesis can be decomposed into a passive componentand an active control component. The active control component is analgebraic function of the user's real-time input F_(s) (i.e., sensedsocket-prosthesis interface forces and moments and sensed groundreaction forces). The controller output is shown as the active torque(τ_(a)) minus the passive torque τ_(p). The controller outputτ_(a)−τ_(p) applied to the prosthetic leg based on dynamics of the legresponds via θ and {dot over (θ)}. The system response, θ and {dot over(θ)}, is fed back to the controller.

Power applied to the prosthesis can be thus commanded directly by theuser through measured interface forces and moments initiated by usermovements. In the absence of these commands from the user, F_(s)=0,τ_(a)=0 and the prosthesis fundamentally (by virtue of the controlstructure) cannot generate power, and thus only exhibits controlledpassive behavior. Due to the decomposition of energetic behaviorsinherent in this control structure, the prosthesis under its own controlcan be generally stable and passive. Unlike known echo controlapproaches, the input can be real-time, based only on the affected leg,and thus the approach can be equally applicable to bilateral andunilateral amputees and can reflect the instantaneous intent of theuser. Additionally, unlike echo control that is based on servocontrol,the prosthesis will exhibit a natural impedance to the user that shouldfeel more like a natural limb. These combined features should result inan active prosthesis that will feel inasmuch as possible like a naturalextension of the user. The structure and properties of both the gaitcontroller and intent recognizer are described below.

As described above, since gait is largely a periodic activity, jointbehavior can be functionally decomposed over a period by decomposing thejoint torque into a passive component and an active component. Thepassive component can comprise a function of angle (i.e., single-valuedand odd), and a function of angular velocity passive (i.e.,single-valued and odd), such as equation 1 described above. The activecomponent can be a function of the user input (i.e., socket interfaceforces). Given a set of data that characterizes a nominal period ofjoint behavior, the passive component can be first extracted from thewhole, since the passive behavior is a subset of the whole (i.e., thepassive component consists of single-valued and odd functions, while theactive has no restrictions in form). The passive component can beextracted by utilizing a least squares minimization to fit a generalizedsingled-valued odd function of angle and angular velocity to the torque.Once the passive component is extracted, the residual torque (i.e., theportion that is not extracted as a passive component), can beconstructed as an algebraic function of the sensed socket interface andground reaction forces (i.e., the direct-acting user input) byincorporating a similar candidate function, but not restricted to be ofpassive form. Finally, superimposing the passive and active componentsprovides a decomposed functional approximation of the original periodjoint torque.

In the embodiment of the intra-modal controller shown in FIG. 12, asupervisory intent recognizer can be added that utilizes the same senseduser inputs (i.e., moments and forces) as the intra-modal/gaitcontroller, but extracts the user's intent based on the characteristicshape of the user input(s) and system response (e.g. F, θ, θ-dot). Basedon the extracted intent, the supervisory intent recognizer modulates thebehavior of the underlying gait controller to smoothly transitionbehavior within a gait (e.g., speed and slope accommodation) and betweengaits (e.g., level walk to stair ascent), thus offering a unifiedcontrol structure within and across all gaits.

Gait intent recognition can be a real-time pattern recognition or signalclassification problem. The signal in this case is generally thecombination of socket interface forces Fs and the dynamic state of theprosthesis, which in one embodiment can be a vector of the knee andankle angles θ for a powered leg prosthesis according to an embodimentof the invention. A variety of methods exist for pattern recognition andsignal classification including nearest neighbor algorithms, neuralnetworks, fuzzy classifiers, linear discriminant analysis, and geneticalgorithms.

As described above, embodiments of the invention include a number ofsensors for providing signals for adjusting operation of a leg and ankleprosthesis. A description of one exemplary arrangement of sensors can bedescribed below with respect to FIGS. 13A, 13B, 14A, and 14B. FIG. 13Ais a side view of powered knee and ankle prosthesis 1300, according toanother embodiment of the invention. FIG. 13B is a front view of poweredknee and ankle prosthesis of FIG. 13A. FIGS. 14A and 14B showperspective and bottom views of an exemplary sagittal moment load cellsuitable for use in the various embodiments of the invention.

Each joint actuation unit, such as knee actuation unit 1302 and ankleactuation unit 1304 in FIG. 13A, can include a uniaxial load cellpositioned in series with the actuation unit for closed loop forcecontrol. Both the knee and ankle joints can incorporate integratedpotentiometers for joint angle position. The ankle actuation unit caninclude a spring 1305, as described above with respect to FIGS. 1A-4.One 3-axis accelerometer can be located on the embedded system 1306 anda second one can located below the ankle joint 1308 on the ankle pivotmember 1310. A strain based sagittal plane moment sensor 1312, such assensor 1400 shown in FIGS. 14A and 14B, can located between the kneejoint 1314 and the socket connector 1316, which measures the momentbetween a socket and the prosthesis. In the various embodiments of theinvention, a sagittal plane moment sensor can be designed to have a lowprofile in order to accommodate longer residual limbs. The sensor canincorporate a full bridge of semiconductor strain gages which measurethe strains generated by the sagittal plane moment. In one embodiment ofthe invention, the sagittal plane moment sensor was calibrated for ameasurement range of 100 Nm. A custom foot 1318 can designed to measurethe ground reaction force components at the ball 1320 of the foot andheel 1322. The foot can include of heel and ball of foot beams, rigidlyattached to a central fixture and arranged as cantilever beams with anarch that allows for the load to be localized at the heel and ball ofthe foot, respectively. Each heel and ball of foot beam can alsoincorporates a full bridge of semiconductor strain gages that measurethe strains resulting from the respective ground contact forces. In oneembodiment of the invention, the heel and ball of foot load sensors werecalibrated for a measurement range of 1000 N. In addition, incorporatingthe ground reaction load cell into the structure of a custom foot caneliminate the added weight of a separate load cell, and also enablesseparate measurement of the heel and ball of foot load. The prostheticfoot can be designed to be housed in a soft prosthetic foot shell (notshown).

The powered prostheses described above contain an embeddedmicrocontroller that allows for either tethered or untethered operation.An exemplary embedded microcontroller system 1500 is shown in the blockdiagram in FIG. 15. The embedded system 1500 consists of signalprocessing, power supply, power electronics, communications andcomputation modules. The system can be powered by a lithium polymerbattery with 29.6 V. The signal electronics require +/−12 V and +3.3 V,which are provided via linear regulators to maintain low noise levels.For efficiency, the battery voltage can be reduced by PWM switchingamplifiers to +/−15 V and +5 V prior to using the linear regulators. Thepower can be disconnected via a microcontroller that controls a solidstate relay. The power status can be indicated by LED status indicatorscontrolled also by the microcontroller.

The analog sensor signals acquired by the embedded system include theprosthesis sensors signals (five strain gage signals and twopotentiometer signals), analog reference signals from the laptopcomputer used for tethered operation, and signals measured on the boardincluding battery current and voltage, knee and ankle servo amplifiercurrents and two 3-axis accelerometers. The prosthesis sensor signalsare conditioned using input instrumentation amplifiers. The battery,knee motor and ankle motor currents are measured by current senseresistors and current sensing amplifiers. The signals are filtered witha first-order RC filter and buffered with high slew rate operationalamplifiers before the analog to digital conversion stage. Analog todigital conversion can be accomplished by two 8-channel analog todigital convertors. The analog to digital conversion data can betransferred to the microcontroller via serial peripheral interface (SPI)bus.

The main computational element of the embedded system can be a 32-bitmicrocontroller. In the untethered operation state, the microcontrollerperforms the servo and activity controllers of the prosthesis and datalogging at each sample time. In addition to untethered operation, theprosthesis can also be controlled via a tether by a laptop computerrunning MATLAB Simulink RealTime Workshop. In the tethered operationstate, the microcontroller drives the servo amplifiers based on analogreference signals from the laptop computer. A memory card can be usedfor logging time-stamped data acquired from the sensors and recordinginternal controller information. The memory chip can be interfaced tothe computer via wireless USB protocol. The microcontroller sends PWMreference signals to two four quadrant brushless DC motor drivers withregenerative capabilities in the second and forth quadrants of thevelocity/torque curve.

As noted above with respect to FIG. 9, additional controls can beprovided for operating the prosthesis when going from a sitting to astanding position or vice versa. This can be implemented via the use ofa sitting mode controller implemented in the microcontroller. Operationof the sitting mode controller consists of four phases that are outlinedin the general control state chart shown in FIG. 16. As shown in FIG.16, two phases are primary sitting phases, weight bearing and non-weightbearing. The other two phases encompass the transition phases, pre-standand pre-sit, for standing up and sitting down, respectively. Weightbearing and non-weight bearing are the primary sitting phases thatswitch the knee and ankle joints between high and low impedances,respectively. The transition phases, pre-stand and pre-sit, modulate thestiffness of the knee as a function of knee angle, as shown in FIG. 17,to assist the user in standing up and sitting down. FIG. 17 shows kneeangle modulated knee stiffness during pre-stand (solid line) and pre-sit(dashed line) phases.

The modulation allows for smoother transitions near the seated position.The ankle joint can be slightly dorsiflexed with moderate stiffnessduring the standing up and sitting down phases. Switching between thefour sitting phases occurs when sensor thresholds are exceeded, asdepicted FIG. 16. The parameters of the impedance based controllers aretuned using a combination of feedback from the user and joint angle,torque and power data from the prosthesis.

In the various embodiments of the invention, actuation for a prosthesiscan be provided by two motor-driven ball screw assemblies that drive theknee and ankle joints, respectively, through a slider-crank linkage. Theprosthesis can be capable of 120° of flexion at the knee and 45° ofplanterflexion and 20° of dorsiflexion at the ankle. In one embodiment,each actuation unit consists of a DC motor (such as a Maxon EC30Powermax) connected to a 12 mm diameter ball screw with 2 mm pitch, viahelical shaft couplings. An exemplary ankle actuation unit additionallyincorporates a 302 stainless steel spring (51 mm free length and 35 mmouter diameter), with 3 active coils and a stiffness of 385 N/cm inparallel with the ball screw.

As described above with respect to FIGS. 1A-4, the purpose of the springcan be to bias the motor's axial force output toward ankleplantarflexion, and to supplement power output during ankle push off.The stiffness of the spring can be maximized to allow for peak forceoutput without limiting the range of motion at the ankle. The resultingaxial actuation unit's force versus ankle angle plot can be shown inFIG. 18. FIG. 18 is a plot if axial force as a function of ankle angleillustrating spring force, actuator force and total force. FIG. 18graphically demonstrates for fast walking the reduction in linear forceoutput supplied by the motor at the ankle through the addition of thespring. Note that the compression spring does not engage untilapproximately five degrees of ankle plantarflexion. Each actuation unitcan include a uniaxial load cell (such as Measurement SpecialtiesELPF-500L), positioned in series with the actuation unit for closed loopforce control of the motor/ballscrew unit. Both the knee and anklejoints can incorporate bronze bearings and, for joint angle measurement,integrated precision potentiometers (such as an ALPS RDC503013). Astrain based sagittal plane moment sensor, as previously described withrespect to FIGS. 14A and 14B can be located between the knee joint andthe socket connector, which measures the moment between the socket andprosthesis. The ankle joint connects to a foot, which incorporatesstrain gages to measure the ground reaction forces on the ball of thefoot and on the heel. The central hollow structure houses alithium-polymer battery and provides an attachment point for theembedded system hardware. To better fit with an anthropomorphicenvelope, the ankle joint can be placed slightly anterior to thecenterline of the central structure. This gives the prosthesis theillusion of flexion when the amputee can be standing vertically with theknee fully extended.

The length of the shank segment can be varied by changing the length ofthree components; the lower shank extension, the spring pull-down, andthe coupler between the ball nut and ankle. Additional adjustability canbe provided by the pyramid connector that can be integrated into thesagittal moment load cell for coupling the prosthesis to the socket (asis standard in commercial transfemoral prostheses).

Passive joint torque, τ_(p), can be defined as the part of the jointtorque, τ, which can be represented using spring and dashpotconstitutional relationships (passive impedance behavior). The systemcan only store or dissipate energy due to this component. The activepart can be interpreted as the part which supplies energy to the systemand the active joint torque can be defined as τ_(a)=τ−τ_(p). This activepart can be represented as an algebraic function of the user input viathe mechanical sensory interface (i.e socket interface forces andtorques).

Gait is considered a mainly periodic phenomena with the periodscorresponding to the strides. Hence, the decomposition of a stride willgive the required active and passive torque mappings for a specificactivity mode. In general, the joint behavior exhibits varying activeand passive behavior in each stride. Therefore, segmenting of the stridein several parts can be necessary. In this case, decomposition of thetorque over the entire stride period requires the decomposition of thedifferent segments and piecewise reconstruction of the entire segmentperiod. In order to maintain passive behavior, however, the segmentscannot be divided arbitrarily, but rather can only be segmented when thestored energy in the passive elastic element is zero. This requires thatthe phase space can only be segmented when the joint angle begins andends at the same value. FIG. 19 shows the phase portrait of normal speedwalking and the four different stride segments, S₁, S₂, S₃ and S₄. Thus,the entire decomposition process consists of first appropriatesegmentation of the joint behavior, followed by the decomposition ofeach segment into its fundamental passive and active components.

The decomposition of each segment shown in FIG. 19 can be converted toan optimization problem. In each segment of the stride, 2n data pointsare selected by sampling the angular position in equal intervals betweenits minimum and maximum and selecting the corresponding positive andnegative angular velocities. In this work, the number of angularposition samples for each segment, n can be set to be 100. Theconstrained least squares optimization problem given in Equation 2 belowcan be constructed and solved.

$\begin{matrix}{{\min\limits_{x}{\frac{1}{2}{{{Cx} - d}}_{2}^{2}\mspace{14mu}{s.t.\mspace{14mu} 0}}} \leq x} & (2)\end{matrix}$where C, x and d are defined in Equations 3, 4, and 5 below,respectively. The indexing of the joint angular position, angularvelocity and moment samples are explained via the sketch in FIG. 20.FIG. 20 shows a selection and indexing of data samples from a firstsegment.

$\begin{matrix}{{{C_{4{nx}\; 3n} = \begin{bmatrix}C_{1} & C_{2} & C_{3}\end{bmatrix}^{T}}{C_{1} = \lbrack {\begin{matrix}{{diag}( {\begin{bmatrix}\theta_{1} \\\theta_{2} \\\theta_{3} \\\vdots \\\theta_{n}\end{bmatrix}_{{nx}\; 1} - \alpha} )} \\{{diag}( {\begin{bmatrix}\begin{matrix}\begin{matrix}\theta_{n} \\\theta_{n - 1}\end{matrix} \\\vdots\end{matrix} \\\theta_{1}\end{bmatrix}_{{nx}\; 1} - \alpha} )}\end{matrix}{{diag}( \begin{bmatrix}{\overset{.}{\theta}}_{1} \\{\overset{.}{\theta}}_{2} \\\vdots \\\vdots \\\vdots \\{\overset{.}{\theta}}_{n}\end{bmatrix}_{2n\; x\; 1} )}} \rbrack}{C_{2} = \lbrack {\begin{matrix}C_{21} \\C_{22}\end{matrix}C_{23\;}} \rbrack_{{2n} - {1x\; 3n}}}C_{21} = \begin{bmatrix}\theta_{1} & {- \theta_{2}} & 0 & \ldots & 0 \\0 & \ddots & \ddots & \ddots & \vdots \\\vdots & \ddots & \theta_{n - 1} & \theta_{n} & 0 \\0 & \ldots & 0 & 0 & 0\end{bmatrix}_{nxn}}{C_{22} = \begin{bmatrix}\theta_{n} & {- \theta_{2}} & 0 & \ldots & 0 \\0 & \ddots & \ddots & \ddots & \vdots \\\vdots & \ddots & \theta_{3} & {- \theta_{2}} & 0 \\0 & \ldots & \theta_{3} & {- \theta_{2}} & 0\end{bmatrix}_{n - {1{xn}}}}{C_{23} = \begin{bmatrix}{\overset{.}{\theta}}_{n} & {- {\overset{.}{\theta}}_{n - 1}} & 0 & \ldots & 0 \\0 & \ddots & \ddots & \ddots & \vdots \\\vdots & \ddots & {\overset{.}{\theta}}_{{2n} - 2} & {- {\overset{.}{\theta}}_{{2n} - 1}} & 0 \\0 & \ldots & 0 & {\overset{.}{\theta}}_{{2n} - 1} & {- {\overset{.}{\theta}}_{2n}}\end{bmatrix}_{{2n} - {1x\; 2n}}}{C_{3} = \begin{bmatrix}\beta & \beta & \ldots & \ldots & \ldots & \beta & \beta\end{bmatrix}_{1x\; 3n}}} & (3) \\{x_{3{nx}\; 1} = \begin{bmatrix}k_{1} \\k_{2} \\\vdots \\k_{n - 1} \\k_{n} \\b_{1} \\b_{2} \\\vdots \\b_{{2n} - 1} \\b_{2n}\end{bmatrix}} & (4) \\{d_{4{nx}\; 1} = \begin{bmatrix}\tau_{1} \\\tau_{2} \\\vdots \\\tau_{{2n} - 1} \\\tau_{2n} \\{\tau_{1} - \tau_{2}} \\{\tau_{2} - \tau_{3}} \\\vdots \\{\tau_{{2n} - 1} - \tau_{2n}} \\0\end{bmatrix}} & (5)\end{matrix}$

The matrix C consists of three sub-matrices, C₁, C₂ and C₃. C₁ can bethe main part responsible for the fitting of the spring and dashpotconstants, k and b. C₂ bounds the rate of change of the passive jointtorque and ensures smoothness in the resulting passive joint torque, andC₃ is basically a row of penalty constants, β, which penalizes largevalues of the spring and dashpot constants and thus limits themagnitudes of both. In this work, β is set to 0.1.

The origin of each virtual spring can be also added to the optimizationproblem formulation as a parameter in order to obtain a tighter passivetorque fit. Therefore, the optimization problem given by (3) can besolved iteratively for a range of values of spring origin constant, α.The solution with the least error norm can be selected as the optimalsolution.

The result of the above stated constrained optimization problem forsegment 1 can be shown in the plots shown in FIGS. 21A, 21B, and 21C.These plots are the output of the decomposition for s₁ in FIG. 19showing the spring constant (FIG. 21A), dashpot constants (FIG. 21B),and the active and passive knee torques (FIG. 21C; Spring origin, α. is23 degrees).

As can be seen from FIGS. 21A, 21B, and 21C, the decomposed passive partcan be very similar to the joint torque, and thus it can be stated thatthe behavior of the joint can be mainly passive. The result of thedecomposition for the segment S_(i) can be stored in R_(i) of the formgiven in Equation 6.R _(i)=[θ{dot over (θ)}τ_(pas) F _(S1) F _(S2)τ_(act)]_(2n×6)  (6)where τ_(pas)=C₁x.

The procedure presented above decomposes the joint torques into activeand passive parts. The joint torque references for the control of theprosthesis are generated by combining this active and passive torques.There are two major challenges to be solved. Firstly, the correct motionsegment must be selected. Secondly, after the motion segment is selectedat each sampling instant a new joint torque reference can be generatedusing the discrete mappings for the active and passive torque parts.

A switching system modeling approach incorporating both discrete andcontinuous states can be used for the reconstruction of the torquereference signal. The state chart shown in FIG. 22. will govern thediscrete dynamics of the controller. Since the sequence of the segmentscan be ordered (i.e., the direction of the motion for a specific gaitphase does not change), each segment can transition only to the nextone, where the transition guard function can be written as a inequalityin terms of θ and {dot over (θ)}. The transitions between segments takeno time and the dynamics of the controller are governed by the{ƒ_(pi)(θ,{dot over (θ)});ƒ_(a) _(i) (F_(S))} pair at each samplinginstant. The joint reference torque isτ_(ref)=τ_(a)+τ_(p)=ƒ_(p) _(i) (θ,{dot over (θ)})+ƒ_(a) _(i) (F_(S))  (7)

The decomposition algorithm presented above gives the result matrix, R,for each segment. The discrete data in R can be used to construct thejoint torque reference for the continuous measurements of another trialin the same gait phase. At each sampling instant of the algorithm, themeasurement vector m=[θ_(m),{dot over (θ)}_(m),F_(S1_m),F_(S2_m)]^(T)can be acquired. For the reconstruction of the passive knee torque part,the Euclidian error norm between the [θ_(m) {dot over (θ)}_(m)]^(T) andthe angular position and velocities of all the samples in that segment[θ_(i) {dot over (θ)}_(i)]^(T) can be calculated as shown in Equation 8and stored in the vector e.e _(i)=√{square root over ((θ_(m)−θ_(i))²+({dot over (θ)}_(m)−{dot over(θ)}_(i))²)}  (8)Then two elements of this vector with the least error norm are found andthe passive knee torque reference can be found as a weighted linearcombination of the passive knee torques corresponding to these points.The reconstruction of the active knee torque part is similar where only{θ,{dot over (θ)},τ_(pas)} is exchanged with {F_(S1),F_(S2),τ_(act)}.

The supervisory controller (intent recognizer) switches among differentunderlying intramodal controllers depending on the activity mode theuser imposes on the prosthesis. The intent recognizer consists of threeparts: activity mode recognizer, cadence estimator and the slopeestimator.

The activity mode recognizer detects the activity mode of the prosthesis(standing, walking, sitting, stair ascent or stair descent, etc. . . .). This can be accomplished by comparing the features which aregenerated in real time to a feature database using some machine learningand/or pattern recognition methods. The present implementation of thegait mode recognizer, which recognizes standing and walking modes, isdescribed below.

Firstly, a database which contains all the possible activity modes(standing and walking in this case) can be generated by makingexperimental trials. In the experimental trials, the user can be askedto walk or stand in different controller modes for 50 second longtrials. The socket sagittal moment above the knee joint, foot heel load,foot ball load, knee angle, knee velocity, ankle angle and anklevelocity are recorded with 1 ms sampling period. It should be noted thatother sensor signals such as accelerations and electromyographymeasurements from the residual limb can be added to the list of thesignals used for intent recognition. For example, from the recordedexperimental trials, 10000 random frames (5000 standing and 5000walking) of 100 samples length are generated for all the seven recordedsignals. The mean and the standard deviation of each frame are computed.The mean and standard deviation of signals are selected as the featuressince minimal computation can be required to obtain them. A databasecontaining 10000 samples with 14 features (mean and standard deviationof the seven signals) belonging to two classes (standing and walking)can be generated. After the database is generated, the dimension of thedatabase can be reduced from 14 to three using principal componentanalysis (PCA). Dimension reduction can be necessary because patternrecognition for high dimensional datasets can be computationallyintensive for real-time applications. After dimension reduction step,the standing and walking data can be modeled with Gaussian mixturemodels. Gaussian mixture models represent a probability distribution asa sum of several normal Gaussian distributions. The order of theGaussian mixture model for each mode can be determined according to theMinimum Description Length Criteria.

As described above, the database generation, dimension reduction and theGaussian mixture modeling are explained. For real-time decision making,overlapping frames of 100 samples can be generated at each 10 msinterval. 14 features described above are extracted from these framesand the PCA dimension reduction can be applied to these features to geta reduced three dimensional feature vector. The reduced dimensionfeatures can be fed to the Gaussian mixture models for standing andwalking and the probability of the sample vector being standing orwalking can be computed. The mode with the greater probability isselected as the instantaneous activity mode. Since one decision mightgive wrong results in some cases due to noise, disturbance, etc. . . . ,a voting scheme can be used to enhance the results. In the votingscheme, the controller activity mode is switched if and only if morethan 90 percent of the instantaneous activity mode decisions among thelast 40 decisions are a specific activity mode. Once a new activity modeis selected by the voting scheme, the underlying activity controller canbe switched to the corresponding mode.

Such an activity mode recognizer is provided by way of illustration andnot as a limitation. In the various embodiments of the invention, one ormore parts of the algorithm might be modified. For example, in someembodiments, different features such as mean, max, kurtosis, median, ARcoefficients, wavelet based features, frequency spectrum based featuresof the frame might be generated. Additionally, different dimensionreduction techniques such as linear discriminant analysis, independentcomponent analysis might be employed. Furthermore, differentclassification methods such as artificial neural networks, supportvector machines, decision trees, hidden Markov models might be used.

Cadence estimation is accomplished by observing peak amplitudes incharacteristic signal data and then measuring the time betweensuccessive peaks. Since walking is a cyclic activity each of the sensorsignals will be periodic of cadence. The most relevant sensor signalswill contain only one characteristic amplitude peak per stride such asfoot heel load and the ball of foot load. In the real-timeimplementations, cadence estimation is accomplished by recording thefoot load after heel strike when it exceeds 400 N until the loaddecreases below 350 N. Then, the time of occurrence of the peak load inthis window is found and the previous peak time is subtracted from thenew peak time. This corresponds to stride time and can be converted tocadence (steps/min) by multiplying with 120. Once the cadence isestimated, the intent recognizer selects the corresponding middle layercontroller based on some predefined thresholds as in FIG. 23.

For example, in some embodiments, a 3D accelerometer capable ofmeasuring ±3 g accelerations is embedded into the ankle joint couplerwhere the prosthetic foot is connected. An exemplary arrangement of sucha system is shown by the schematic in FIG. 24. The accelerometermeasurements are used to estimate the ground slope. In order to estimatethe ground slope, the accelerometer data in tangential direction isused. Assuming the foot is flat on the ground, the ground slope angle,θ_(s), can be calculated as in equation (9) below.

$\begin{matrix}{\theta_{s} = {\sin^{- 1}( \frac{a_{t}}{g} )}} & (9)\end{matrix}$In Eqn. 9, g is the gravitational constant. In order to find the groundslope estimate, {circumflex over (θ)}_(s), the accelerometer data shouldbe collected while the foot is flat on the ground as determined by theheel and ball of the foot load sensors. While the foot is flat on theground, equation (1) is computed for the frame of the collected data andthe mean of this frame is outputted as the ground slope estimate,{circumflex over (θ)}_(s). Once the slope is estimated, the intentrecognizer selects the corresponding middle layer controller based onsome predefined thresholds. An exemplary state chart for such an intentrecognizer is shown in FIG. 25.

Rather than a ballscrew and slider crank embodiment for the transmissionof torque from a motor to the ankle and/or knee units, in someembodiments of the invention, the prosthesis can incorporate a frictionand cable drive transmission embodiment. FIGS. 26A and 26B show frontand back views of an exemplary embodiment of a friction drivetransmission 2600 in accordance with an embodiment of the invention. Asshown in FIGS. 26A and 26B, the shaft 2602 of an electric motor 2604 ispreloaded against a first stage in a housing 2606, such as a largerdiameter cylinder or friction drive gear 2608, which creates sufficientfriction to transmit torque without slip. The shaft 2602 can use one ormore friction rollers 2610 to transmit the torque. The first stage ofthe friction drive can also be supplemented with a second stage. Thefriction drive gear 2608 drives a smooth pinion 2612 directly, which ispreloaded against a larger diameter cylinder or cable gear output 2614in the housing 2606, which in turn transmits torque directly to the kneeor ankle joint.

In addition to, or rather than a friction drive, the first or secondstage of the transmission can alternatively be embodied by a cable drivetransmission, in which a cable is wrapped around the circumference of alarger diameter cylinder, such as friction drive gear 2608, and alsoaround the circumference of a smaller diameter cylinder, such as pinion2612. In such embodiments, the cable is affixed to the friction drivegear 2608, and is pretensioned, using a tensioning screw 2616 or similarmeans, around both the drive gear 2608 and pinion 2612, such thatfriction between cable and pinion 2612 enables the transmission oftorque from between the pinion 2612 and drive gear 2608. In oneembodiment of a combined friction drive/cable drive transmission can beused, in which a first stage of the transmission (i.e., the frictiondrive gear 2608 connected directly to the electric motor 2604) is of thefriction drive type, while the second stage of the transmission (i.e.,the cable gear output 2614 connected directly to the knee or anklejoint) is of the cable drive type.

Rather than the ballscrew and slider crank or the friction drive andcable drive embodiments for the transmission of torque from a motor tothe ankle and/or knee units, in some embodiments of the invention, theprosthesis can incorporate a chain drive or a belt drive transmissionembodiment for implementing one or more stages of a transmission.

Advantages of a belt or chain drive approach over the ballscrewapproaches described above include the ability to provide a fullyenclosable/sealable (without need for a bellows-type cover) powered legdevice. This facilitates component immersion in lubricating environment,and well as facilitating isolation from dirt, water, and other debris.As a result, this can extend the lifetime of transmission components.Another advantage of such a configuration is that it enables a greaterrange of motion of joint actuation, as opposed to a slider-crankmechanism (as used in a ballscrew configuration), which is generallylimited. Further, the belt or chain drive approach also allows thedevice to maintain a constant transmission ratio throughout range ofmotion, which is not generally possible in the slider-crank mechanismtypically used in a ballscrew configuration. Additionally, advantages ofa belt or chain drive approach is that it maintains constant mechanismgeometry throughout range of motion, belt and chain drive components aretypically less expensive than ballscrew components, and belt and chaindrive systems are typically characterized by lower audible noise thanballscrew configurations.

FIG. 27 shows an exemplary embodiment of a belt drive transmission 2700in accordance with an embodiment of the invention. As shown in FIG. 27,a stage of the transmission 2700 can be embodied as a belt drivetransmission, in which a belt 2702 is wrapped around the circumferenceof a larger diameter shaft, such as a first belt gear or pulley 2704,and also around the circumference of a smaller diameter shaft, such assecond belt gear or pulley 2706. In such embodiments, the belt 2702 canbe tensioned, using a tensioning device 2708. In one embodiment, thetensioning device 2708 can consist of a swing arm 2710, an additionalpulley 2712 attached to the end of swing arm 2710, and tensioning screw2714 for adjusting the swing arm 2710 to bias the additional pulley 2712against the belt 2702, such that friction between the belt 2702 and beltgears 2704 and 2706 enables the transmission of torque from betweensecond belt gear 2706 and first belt gear 2704. However, any other typeof tensioning device can be used in the various embodiments to tensionthe belt 2702. For example, in some embodiments, the tensioning device2708 can be a spring loaded device to automatically bias a pulley 2706or other object against belt 2702 to cause the necessary tension.

It is worth noting that although transmission 2700 is illustrated interms of a V-belt embodiment, the invention is not limited in thisregard and can be used with any type of belts. For example, the belt2702 can also be embodied as a flat belt, a round belt, a multi-groovebelt, a ribbed belt, and a toothed or cog belt, to name a few. Further,the belt gears 2704 and 2706 can be configured in accordance with thetype of belt being used.

In some embodiments, rather than utilizing a belt-based drive, achain-based drive can be provided. The configuration in such embodimentscan be substantially similar to that shown in FIG. 27. That is, a chaincan be provided in place of belt 2702 and gears 2704 and 2706 can beembodied as sprockets compatible with the chain. In such embodiments,the tensioning device 2708 described above can still be utilized tomaintain proper tension of the chain to enable the transmission oftorque from between sprockets in the transmission.

In some embodiments, instead of utilizing a tensioning device asdescribed above with FIG. 27, a pulley or sprocket can be configuredwith an eccentric mount. That is, configuring at least one of the drivegears in the transmission to allow an adjustment of its position. Thisis illustrated below with respect to FIGS. 28A-28D.

FIGS. 28A and 28B show side views of first and second positions,respectively, achievable for an exemplary embodiment of a chain drivetransmission 2800 including an eccentric mount in accordance with anembodiment of the invention. Similar to the transmission described abovewith respect to FIG. 27, transmission 2800 includes a first shaft 2802with first drive gears or sprockets 2804 and a second shaft 2806 withsecond drive gears or sprockets 2808 which can be coupled together viachains 2810 to transmit torques between sprockets 2804 and sprockets2808. Although FIGS. 28A and 28B show that the transmission of torquebetween sprockets 2804 and sprockets 2808 is performed using two sets ofsprockets (and thus using two chains), the embodiments are not limitedin this regard. Rather, any number of chains can be used in the variousembodiments.

As shown in FIGS. 28A and 28B, the first shaft 2802 is shown asincluding an additional sprocket 2812 for driving first shaft 2802. Sucha configuration can be used when multiple drive stages are provided.However, the various embodiments are not limited in this regard.

In transmission 2800, the first shaft 2802 is configured to beeccentric. That is, the position of the first shaft 2802 is adjustablerelative to the position of the second shaft 2806 so as to adjust thelateral separation between the shafts (i.e., to provide d_(A)≠d_(B)).Accordingly, this also provides a means to adjust the tension in a chain(or a belt) between the first shaft 2802 and the second shaft 2806. Toprovide the eccentric mount, the first shaft 2802 can be mounted in aleg device to an adjustable bearing mount 2814. The operation andconfiguration of an exemplary embodiment of the adjustable bearing mount2814 is illustrated with respect to FIG. 29.

FIG. 29 illustrates schematically the components for the adjustablebearing mount 2812. As shown in FIG. 29, the adjustable bearing mount2814 can include a top plate 2902 to which first shaft 2802 is attached,a bottom plate 2904, bearings 2906 between the top plate 2902 and thebottom plate 2904, and fasteners 2908. These components of theadjustable bearing mount 2814 can be disposed within an enclosure 2910.

In FIG. 29, the fasteners 2908 are shown as screws or bolts. However,the various embodiments are not limited to any particular bearing typeor design of screws or bolts and other bearing types or designs can beused without limitation. Further, the various embodiments are notlimited to screws or bolts and any other type of removable fastener canbe used without limitation. Additionally, FIG. 29 shows bearings 2906 asa collection of ball bearings disposed between plates 2902 and 2904.However, the various embodiments are not limited to any particularbearing type or design and other bearing types or designs can be usedwithout limitation.

In operation, the enclosure 2910 can be configured such that whenfasteners 2908 are loosened or removed, the bearings allow the top plate2902 can be repositioned relative to the bottom plate 2904 via bearings2906. Thus, when fasteners 2908 are replaced and tightened, the plates2902 and 2904 are biased against bearings 2906 to prevent further motionof the top plate 2902 relative to the bottom plate 2904.

Such a configuration allows adjustment of the position of first shaft2802. For example, this can allow the first shaft 2802 to transitionbetween a first position, as shown in FIG. 28A, in which a chain or belt2810 with reduced tension is provided, due to a reduced distance (d_(A))between first shaft 2802 and second shaft 2806, to a second position, asshown in FIG. 28B, in which a chain or belt 2814 with increased tensionis provided, due to an increased distance (d_(B)) between first shaft2802 and second shaft 2806. However, the various embodiments are notlimited to solely first and second positions. Rather, in the variousembodiments, the adjustable bear mount 2812 can be configured to allow avariety of positions for the first shaft 2806 relative to the secondshaft 2806.

An exemplary configuration of a powered leg prosthesis 3000 inaccordance with the discussion above is illustrated schematically inFIG. 30. As shown in FIG. 30, the powered leg prosthesis 3000 includes ashank 3002 with a powered knee joint 3004 and a powered ankle joint3006. The powered knee joint 3004 includes a socket interface 3008 forattaching a socket 3010 or other device for attachment of the poweredleg prosthesis 3000 to an amputee. The powered ankle joint 3006 can havea foot portion 3012 attached thereto.

The shank 3002 can consist of a single, discrete unit. However, in someembodiments, the shank can include an upper portion 3014 and a lowerportion 3016. Such a configuration allows the insertion of at least oneextension unit 3018 to allow the length of the shank 3002 to becustomized for the amputee.

Within each of the upper portion 3014 and the lower portion 3016, a beltor chain drive system can be implemented, as described above withrespect to FIGS. 27-29. For example, as shown in FIG. 30, the upperportion 3014 can include a first motor 3022, a first upper drive stage3024, and a second upper drive stage 3026 for providing power at thepowered knee joint 3004. Similarly, the lower portion 3016 can include asecond motor 3028, a first upper drive stage 3030, and a second upperdrive stage 3032 for providing power at the powered ankle joint 3006.Each stage can consist of the belt or chain drive stage. Additionally,each stage can be configured to include an eccentric mount, such asmounts 3034 and 3036, to adjust tension in the upper portion 3014 andlower portion 3016 respectively.

In addition to the components described above, the powered prostheticleg 3000 can include other components not illustrated in FIG. 30 forpurposes of clarity. For example, the powered prosthetic leg can includea control system or device, as previously described, and one or moresensors throughout the powered prosthetic leg, also as previouslydescribed. Thus control of the powered prosthetic leg 3000 can occurinsubstantially the same manner as described above.

Running Controller

Now that the configuration and operation of an exemplary lower limbdevice (a single powered prosthesis) has been described, the disclosurenow turns to a description of the running controller in accordance withthe various embodiments. It should be noted that the examples andresults are presented below are provided solely for illustrating thevarious embodiments and are not intended to limit the variousembodiments in any way.

I. Running Controller Considerations

A bipedal running gait is fundamentally distinct from a bipedal walkinggait by the fact that, in the former, each foot is in the air more oftenthan on the ground, while in walking, each foot is on the ground moreoften than in the air. In particular, the amount of time each foot is onthe ground is referred to as the “stance” phase of gait. In walking,each leg remains in the stance phase for approximately 60% of the time,while in a bipedal running gait, each leg is in stance phaseapproximately 40% of the time. Assuming the two legs in a bipedal gaitare completely out of phase, if each is in stance 40% of the time, thenit follows that both must be in the air approximately 20% of the time,which can be referred to as the flight phase of running. In contrast, inwalking, both feet are on the ground 20% of the time, which is referredto as the double-support phase of walking.

In order to sustain a flight phase in running, an amount of verticalpropulsive energy must be produced by each leg at least equal to theamount of energy absorbed by the respective leg during the stance phase.Specifically, the stance phase of a running gait consists essentially oftwo phases: an absorption phase, which lasts roughly for the first halfof the stance phase, and a propulsion phase, which lasts roughly for thesecond half of the stance phase. In the absorption phase, the knee andankle joints function synchronously to absorb energy as the foot landson the ground (called foot strike). During this absorption phase, thecenter of mass of the body is decelerated as it is lowered from theflight phase of the running gait. The absorption phase is followed bythe propulsion phase, in which the knee and ankle joints synchronouslygenerate power, which accelerates the center of mass of the body upwardto set up the flight phase of the running gait.

In contrast, in a walking gait, the knee and ankle joints are notcharacterized by synchronous power absorption and generation. Inparticular, the ankle is characterized by absorption of a small amountof energy (relative to running) during most of the stance phase,followed by a generation of energy (a propulsion phase) similar to arunning gait. The knee joint, however, behaves essentially passively.Specifically, the knee absorbs energy immediately following heel strike,but is characterized by energy absorption rather than generation duringthe latter period of the stance phase as well. In particular, unlike inrunning, the knee absorbs power during the portion of stance in whichthe ankle is generating power. As such, unlike in running, the kneecannot be characterized in the stance phase of walking as consisting ofsubstantially an absorption and propulsion phase. Unlike in running, theknee and ankle do not act energetically in synchrony in the stance phaseof walking. And unlike in running, the amount of power generated by theknee in walking is not greater than or essentially equal to the amountof power absorbed. Note that all these fundamental distinctions areaspects of the stance phase.

FIG. 31 shows the (body-mass-normalized) power characteristics of theknee and ankle joints during the stance phase of running for healthysubjects. As shown therein, FIG. 31 clearly shows that running stancephase consists of an absorption phase (negative net power in bothjoints) and a propulsion phase (positive net power in both joints).Further, FIG. 31 also clearly shows that the two joints are essentiallyoperating in synchrony. FIG. 32 shows the (body-mass-normalized) powercharacteristics of the knee and ankle during the stance phase of walkingfor healthy subjects. In contrast to FIG. 31, FIG. 32 clearly shows theabsence of the two phases, particularly for the knee joint, and clearlyindicating the lack of energetic synchrony. Further, FIG. 31 alsoclearly shows that the knee generates net energy during the stance phaseof running, while FIG. 32 clearly shows the knee dissipates net energyduring the stance phase of walking. It should be noted that the swingphase of running and walking (not shown in the figures) are essentiallythe same, wherein the fundamental behavior of both entails the initialflexion and subsequent extension of the knee (although running is moreexaggerated), and entails relatively little ankle movement.

Due to the fundamental behavioral distinctions during the stance phasebetween a bipedal walking and running gait, a controller for a poweredprosthetic, orthotic, or assistive device that generates a running gaitmust be constructed in a fundamentally different manner from acontroller that provides a walking gait. That is, a walking controllercannot produce a running gait, and vice-versa. In particular, assuming arunning controller is implemented in a finite state construction, thestance phase of a running controller for a prosthesis that consists ofat least a knee joint and possibly also an ankle joint must essentiallyconsist of two finite states, one that provides the essential behaviorof the absorption phase of stance and one that provides the essentialbehavior of the propulsion phase of stance. In the absorption phase, theknee and ankle (if included) joints should dissipate power, while in thepropulsion phase, the knee and ankle (if included) should generatepower, roughly equal to or greater than the amount dissipated. Anexemplary method for power dissipation is by use of a passive impedancefunction. That is, configuring the joint to emulate a passive impedancesuch as a stiffness component, damping component, or both. Bydefinition, a passive system, over time, does not generate net(positive) power. The passive impedance function can consist of the sumof a passive stiffness function and a passive damping function. Such apassive stiffness function should relate joint torque to angle in asingle valued, odd algebraic function, while the passive dampingfunction should relate joint torque to angle in an odd algebraicfunction. In this case, the passivity of the functions will guaranteethat energy is absorbed if the joint is returned to an original at restconfiguration. In the propulsion phase, the knee and ankle (if included)joints should generate power according to any of a number of possiblemethods.

In addition to differing states, a running controller must havediffering switches to move between the respective states. In particular,a condition must exist in a running controller to move from theabsorption phase to the propulsion phase of stance. In an exemplarymethod, this switching can be based on the angular velocity of the kneejoint, and in particular, when the angular velocity of the knee jointreaches zero (i.e., the knee has stopped flexing in the absorption phaseof stance). Alternatively, in another exemplary method, the knee andankle joints can switch from the absorption to the propulsion phase whenthe load measured by the leg has reached a maximum. Note that neither ofthese conditions would work effectively in a walking controller, even ifused strictly for the ankle, since the start of ankle push-off inwalking does not correspond to either a joint angular velocity reversalor peak load measurement. An example of an exemplary running controllerfor a powered prosthesis with a knee and ankle joint is shown in FIG.33.

Since an amputee cannot run with a walking controller, nor walk with arunning controller, a running controller should be complemented with amethod to detect a user's intent to switch between walking and running.In the various embodiments, the controller can switch from a walkingcontroller to a running controller if the detected magnitude of load atheel strike is greater than a threshold load. Similarly, the controllercan switch from the running controller back to a walking controllerbased on the length of time the user is in stance phase.

Other methods can also be used to infer intent to switch between walkingand running modes. For example, the intent recognizer can infer intentto transition between the walking controller and the running controllerbased on a measurement of at least a load or acceleration at footstrike. However, the intent recognizer can also infer intent totransition between a walking controller and the running controller basedon estimation of cadence, a measurement of at least stance time, swingtime, or stride time, or a measurement of at least thigh motion. Thethigh motion can be a measurement of thigh angular velocity.

II. Exemplary Running Controller Implementation

As noted above, a controller of a powered prosthesis can be structuredin three levels. The lowest level controls torque at both the knee andankle joints. The middle level controller generates a torque referencefor the lowest level controller. The middle level controller is afinite-state machine, each state defined by passive impedancecharacteristics for both the knee and ankle. Specifically, the required(knee and ankle) joint torques in each state are characterized by a setof impedance parameters corresponding to the model set forth above inequation 1. Transitions between gait modes or states are triggered bycertain biomechanical conditions being met. Further, a separatecontroller exists for each activity implemented in the prosthesis and atany given time during operation the appropriate middle level controlleris selected by the highest level controller.

The various embodiments of the invention provide a finite-state (middlelevel) controller for running gait that can be integrated into theaforementioned high level controller. Gait modes were determined by aniterative least squares regression application of (1) to a set ofrunning gait data, intended to specify the smallest number of (stable)gait modes which sufficiently modeled healthy running. This model hasfive distinct gait modes (one mode is divided into two submodes) andcorresponding sets of parameters. Within this controller, running gaitis divided into: landing/absorption (Mode 0), propulsion (Mode 1), swingflexion (Mode 2), and swing extension (mode 3), where the propulsionmode can consist of two submodes, push-off (Mode 1A) and toe-off (Mode1B). These modes and their transition conditions are depicted by therunning controller in FIG. 34, which is a specific implementation of therunning controller in FIG. 33. However, it should be noted that theconfiguration in FIG. 34 is presented by way of example and not by wayof limitation. That is, other implementations of the running controllerof FIG. 33 may have more or less features than the configuration of FIG.34.

As noted above, Modes 0 and 1 are stance modes for absorption andpropulsion, respectively, which in healthy running biomechanics, shouldcomprise less than 50% of a stride. Modes 2 and 3 are swing modes, whichin healthy running biomechanics comprise greater than 50% of a stride.While in Mode 0, both the knee and ankle have a relatively highstiffness. The knee flexes in a controlled manner, providing shockabsorption and bearing the user's weight. The ankle initiallyplantarflexes in order to reach a flat-foot state and then dorsiflexesas the user's body center passes over the foot. Both the powered kneejoint and the powered ankle joint yield under the load while theprosthesis is decelerating. Once the knee reaches peak flexion for thestance phase and naturally begins to extend (inferred by a zero crossingin velocity), indicating a natural transition into a power generationphase, the controller transitions into Mode 1, the propulsion phase.

During Mode 1, the knee and ankle actively extend and plantarflex,respectively, in order to propel the user forward and upward. As notedabove, Mode 1 can have two submodes: a Push-off mode (Mode 1A) and aToe-off mode (Mode 1B). During Mode 1A, the knee and ankle activelyextend and plantarflex. Mode 1A will then transition to Mode 1B as soonas the knee reaches peak stance knee extension (a zero crossing in kneevelocity), indicating a natural transition into the swing modecharacterized by knee flexion. During Mode 1B, the knee begins to flexas the ankle continues to plantarflex, which assists in flexion of theknee. Note that mode 1B is very brief and typically coincides withconditions facilitating near immediate transition into Mode 2.Therefore, push-off (Mode 1A) begins with landing and absorption andends when the knee reaches peak stance knee extension. Toe-off (Mode 1B)begins with peak stance knee extension and ends when the foot is off theground.

Upon detecting that the foot is off the ground—for example, once sensorson the prosthesis indicate that the load on the prosthesis has fallenbelow a threshold—Mode 1 will transition to Mode 2. During Mode 2, theknee flexes, and the ankle returns to a slightly dorsiflexed state inorder to prepare for the next heel strike. Mode 2 will then transitionto Mode 3 as soon as the knee naturally begins to extend (a zerocrossing in knee angular velocity), indicating a natural transition intoknee extension. During Mode 3, the knee further extends, preparing forheel strike. Once the heel strike is detected, Mode 3 can transitionback to Mode 0. For example, once sensors on the prosthesis indicatethat the load on the prosthesis is above a load threshold similar tothat employed in the transition from Mode 1 to Mode 2. Further, as shownin the state flow diagram in FIG. 33, if during any aerial mode (modes 2and 3, i.e., where the foot is off the ground) a load is detected (i.e.,foot strike detected), the controller immediately transitions to thelanding or absorption mode (Mode 0).

As noted above, the running controller of FIG. 34 can be complementedwith a method to detect a user's intent (e.g., within the intentrecognizer discussed above) to switch between walking and running. Forexample, the intent recognizer can switch from the walking controller tothe running controller if the detected magnitude of load at foot strikeis greater than a threshold which is greater than the load seen duringheel strike in walking, that is, detecting an attempt at a runningstride by detecting the typical impact on a limb during running. Notethat this threshold is not intended to distinguish whether the user isplacing weight on the prosthesis, as in the running controller, but todistinguish a load typical of running from a load typical of walking.Similarly, the controller can switch from the running controller back toa walking controller based on the length of time the user is in stancephase or swing phase, or the duration of a stride. For example, if theuser remains at Mode 0 and/or Mode 1 for an extended period of time orif the user has a significantly longer stride duration than the previousstride.

Other potential switching conditions between walking and runninginclude 1) a threshold of estimated percent of stride (essentially basedon stride time and estimated cadence) at the start (or end) of swing, 2)a threshold for peak thigh absolute angular velocity during swing, 3) athreshold for load at foot strike, 4) a threshold for shank accelerationat foot strike, and 5) some combination of the above.

III. Running Controller Evaluation

Based on the parameters derived from the aforementioned least squaresregression, the controller's basic function was verified by a healthysubject fitted with an able-bodied adapter, immobilizing the user's kneeat roughly 100° of knee flexion. Once this preliminary verification wascomplete, the prosthesis was fitted to a unilateral transfemoralamputee, and the impedance parameters were tuned to suit the gaitbiomechanics of the amputee subject.

A. Evaluation Metrics

The overarching goal of a running controller in accordance with thevarious embodiments is to enable or improve running gait in the user,specifically in situations when it is not feasible for the user to doffhis or her daily use prosthesis and don a running prosthesis. Moreover,the performance objective of the running controller of the variousembodiments is to reproduce, as faithfully as possible, the functionprovided by the intact limb that the prosthesis has replaced. Thus, therunning controller can be evaluated based on its ability to providesagittal plane joint angles representative of healthy running, thepresence of a double float phase, and on the degree of consistency instride-to-stride gait mode transitions.

In order to obtain reference data representative of healthy running,motion capture data was collected on a small set of healthy subjects.For the motion capture study, five healthy subjects—males ages24-26—each ran on a treadmill at a speed of 2.25 ms⁻¹ for two trials,forty-five seconds each. The motion capture was achieved with twelveOptiTrack S250e high speed infrared cameras running at 120 Hz usingARENA motion capture software. Thirty-four reflective markers wereplaced on each subject corresponding to a full skeleton (similar to theHelen Hayes marker set); the software's skeleton solver was used totrack the subject's motion. The data collected in ARENA was subsequentlyprocessed in MATLAB in order to extract lower limb sagittal jointangles. The joint angles were parsed into single strides (twenty stridesper trial) and normalized to a time base of 100%. An offset was appliedto the ankle for each subject based upon the angle of the foot withrespect to the ground plane during a period where the subject's foot wasknown to be flat on the ground. The mean and standard deviation over allstrides were calculated for each joint.

B. Experimental Tuning

The amputee subject who participated in the running controllerevaluation was a 23-year-old male, 4 years post-amputation. Thesubject's amputation was the result of a traumatic injury; his daily useprosthesis is an Otto-Bock CLeg with a Freedom Innovations Renegadefoot.

The middle level running controller impedance parameters were tunedexperimentally on the treadmill during two sessions. The impedanceparameters and mode transition thresholds employed during the initialcontroller verification with a healthy subject were used as a startingpoint for tuning with the amputee subject, with the spring constants inthe stance modes reduced for user comfort. The impedance parameters wereiteratively tuned based upon a combination of knee and ankle jointangles data, qualitative video analysis, and user feedback/comfort. Theexperimentally tuned impedance parameters are shown in Table 1.

TABLE I IMPEDANCE PARAMETERS Knee Ankle     Gait Mode$\quad\begin{matrix}k \\( \frac{Nm}{\deg} )\end{matrix}$ $\quad\begin{matrix}b \\( \frac{Nms}{\deg} )\end{matrix}$     θ_(eq) (deg) $\quad\begin{matrix}k \\( \frac{Nm}{\deg} )\end{matrix}$ $\quad\begin{matrix}b \\( \frac{Nms}{\deg} )\end{matrix}$     θ_(eq) (deg) 0 4.0 0.1 20.0 5.5 0.2 10.0 1 4.5 0.123.0 3.0 0.1 −18.0 2 3.5 0.2 70.0 2.0 0.1 −18.0 3 3.5 0.15 70.0 1.0 0.15.0 4 0.9 0.15 20.0 3.0 0.2 5.0

Following tuning, the controller was evaluated in trials in which theamputee subject ran on a treadmill at 2.25 ms⁻¹ (5.0 mph); the subjectwas allowed to utilize the treadmill's handrails. Note that, for theamputee subject wearing the powered prosthesis with running controller,this treadmill speed corresponded to a cadence of 130 steps per minute.

FIG. 35 depicts six key elements of a stride captured from a video takenduring one trial. FIG. 36 depicts the mode transitions (percent ofstride)±one standard deviation as recorded during the running controllerevaluations. This figure demonstrates the consistency of gait modetransitions within the running controller. One should note that Mode 1B(toe-off) comprises, on average, less than 3% of stride; this mode wasintended to serve as an overlap for Mode 1A in the ankle and Mode 2 inthe knee, allowing the knee to flex while the ankle continues toplantarflex.

FIG. 37 compares sagittal plane knee joint angles (a) and ankle jointangles (b) for several consecutive strides of the amputee subjectrunning on the powered prosthesis to the same angles for healthysubjects (obtained from the aforementioned healthy subject motioncapture study of running, also at 2.25 ms⁻¹). One should first note thatthe standard deviation of the mean for healthy subjects reflects varietyin the running gaits of healthy subjects. While the healthy subjects didexhibit overall uniformity concerning the features of the joint anglecurves (except in the ankle near toe-off), range of motion variedconsiderably between subjects. Concerning the powered prosthesis, thesalient features of the running gait, in both the knee and ankle angles,generally match those of the healthy subjects. That is, relative towalking, the knee and ankle joints both achieve a considerably greaterdegree of flexion and dorsiflexion, respectively, during the stancephase. The most noticeable deviation between the healthy subject dataand the powered prosthesis data is the slight mismatch in knee jointkinematics in the mode transition from 0 to 1. This may indicate theneed for an additional (brief) gait mode which might better transitionbetween 0 and 1 or a slight adjustment in switching conditions.

As previously mentioned, another significant feature of running gait, inaddition to another significant distinction between walking and running,is that the latter has a stance phase that last less than 50% of thestride, which generates a double float phase of gait (as opposed to thedouble support phase that characterizes walking). Specifically, thestance phase of running has been reported to last between 39% and 45% ofthe stride. Mode 3, which indicates toe-off in the powered prosthesisgait cycle (i.e., the termination of the stance phase), begins onaverage at approximately 45% of stride. As such, the powered prosthesisand running controller provides the relative duration of stance andswing phases that characterizes a running gait and distinguishes it froma walking gait. Visual evidence of the double float phase of gait, asprovided by the powered prosthesis, is shown in FIG. 35.

While various embodiments of the present invention have been describedabove, it should be understood that they have been presented by way ofexample only, and not limitation. Numerous changes to the disclosedembodiments can be made in accordance with the disclosure herein withoutdeparting from the spirit or scope of the invention. Thus, the breadthand scope of the present invention should not be limited by any of theabove described embodiments. Rather, the scope of the invention shouldbe defined in accordance with the following claims and theirequivalents.

Although the invention has been illustrated and described with respectto one or more implementations, equivalent alterations and modificationswill occur to others skilled in the art upon the reading andunderstanding of this specification and the annexed drawings. Inaddition, while a particular feature of the invention may have beendisclosed with respect to only one of several implementations, suchfeature may be combined with one or more other features of the otherimplementations as may be desired and advantageous for any given orparticular application.

The terminology used herein is for the purpose of describing particularembodiments only and is not intended to be limiting of the invention. Asused herein, the singular forms “a”, “an” and “the” are intended toinclude the plural forms as well, unless the context clearly indicatesotherwise. Furthermore, to the extent that the terms “including”,“includes”, “having”, “has”, “with”, or variants thereof are used ineither the detailed description and/or the claims, such terms areintended to be inclusive in a manner similar to the term “comprising.”

Unless otherwise defined, all terms (including technical and scientificterms) used herein have the same meaning as commonly understood by oneof ordinary skill in the art to which this invention belongs. It will befurther understood that terms, such as those defined in commonly useddictionaries, should be interpreted as having a meaning that isconsistent with their meaning in the context of the relevant art andwill not be interpreted in an idealized or overly formal sense unlessexpressly so defined herein.

What is claimed is:
 1. A lower limb device comprising: a powered kneejoint comprising a first motor for providing an output torque at thepowered knee joint; a powered ankle joint comprising a second motor forproviding an output torque at the powered ankle joint; at least onesensor for collecting real-time sensor information for the lower limbdevice; at least one processor communicatively coupled to the at leastone sensor and to the powered knee joint; and a computer-readablemedium, having stored thereon instructions for causing the at least oneprocessor to: operate the powered knee joint and the powered ankle jointaccording to a finite state model for a running mode, wherein the finitestate model comprises an absorption phase, a propulsion phase, a swingflexion phase, and a swing extension phase, wherein the absorption phaseis configured to transition to the propulsion phase, wherein thepropulsion phase is configured to transition to the swing flexion phase,wherein the swing flexion phase is configured to transition to the swingextension phase, and wherein the swing extension mode is configured totransition to the absorption phase, wherein transitions occur repeatedlyuntil an activity mode of the lower limb device changes from the runningmode, wherein the absorption phase comprises: transitioning into theabsorption phase by detecting, at the at least one sensor, that a loadon the lower limb device increased above a first threshold; and, whereinthe knee and ankle joints are each configured in the absorption phase tobehave as a combined stiffness and damping, wherein the stiffness anddamping components are configured such that the powered knee and anklesjoints synchronously absorb energy associated with decelerating avertical motion of a body center of mass; wherein the propulsion phasecomprises: transitioning into the propulsion phase by detecting, at theat least one sensor, that the powered knee joint has absorbed the loadon the lower limb device; and, wherein the knee and ankle joints areeach configured in the propulsion phase such that the powered knee andankles joints synchronously generate energy associated with acceleratingthe vertical motion of the body center of mass.
 2. The lower limb deviceof claim 1, wherein a duration of the propulsion phase is roughly equalto a length of the absorption phase.
 3. The lower limb device of claim1, wherein the propulsion phase further comprises a first sub-mode and asecond sub-mode, wherein the first sub-mode comprises active extensionand plantarflexion of both the powered knee joint and the powered anklejoint, and wherein the second sub-mode comprises flexing of the poweredknee joint while the powered ankle joint continues to plantarflex toassist the flexing of the powered knee joint.
 4. The lower limb deviceof claim 3, wherein the lower limb device transitions from the firstsub-mode to the second sub-mode when the powered knee joint reaches apeak stance knee extension.
 5. The lower limb device of claim 1, whereinthe swing flexion phase comprises: transitioning into the swing flexionphase by detecting, at the at least one sensor, a load on the lower limbdevice decreased below a second threshold; flexing the powered kneejoint; and configuring the powered ankle joint to a dorsiflexed state.6. The lower limb device of claim 1, wherein the swing extension phasecomprises: transitioning into the swing extension phase by detecting, atthe at least one sensor, that a velocity of the powered knee joint is ata threshold velocity for extension; further extending the powered kneejoint; detecting, at the at least one sensor, that a load on the lowerlimb device increased above the first threshold; and completing theswing extension mode.
 7. The lower limb device of claim 1, wherein thecomputer-readable medium further comprises instructions for causing theat least one processor to: selecting the running mode for the lower limbdevice based on the real-time sensor information during a walking modeprior to operating the powered knee joint and the powered ankle joint,wherein a transition between the walking mode and the running mode isbased on a measurement of at least one of a load or acceleration at footstrike, a stance time, a swing time, or a stride time.
 8. The lower limbdevice of claim 1, wherein the transitioning into the absorption phasefurther comprises: providing for a plantarflexion of the powered anklejoint under the load; and providing for, based on the real-time sensorinformation, a dorsiflexion of the powered ankle joint as a center ofmass of the load passes over a foot associated with the powered anklejoint.
 9. The lower limb device of claim 1, wherein the transitioninginto the absorption phase further comprises configuring the stiffnessand damping components for the powered knee joint to cause a flexion ofthe powered knee joint under the load.
 10. The lower limb device ofclaim 1, wherein the transitioning of the propulsion phase furthercomprises detecting, at the at least one sensor, that a velocity of thepowered knee joint is at a threshold velocity for flexion.
 11. The lowerlimb device of claim 1, wherein operating the powered knee joint and thepowered ankle joint further comprises: selecting an impedance parametervalue and an equilibrium angle parameter value for each of the poweredknee joint and the powered ankle joint according to a current phase ofthe finite state model, wherein the impedance parameter valuecharacterizes a resistance to change in a joint angle and wherein theequilibrium angle parameter value indicates the joint angle required inthe absence of the load, and generating an output torque of the poweredknee joint and an output torque of the powered ankle joint according toa pre-defined relationship between a joint torque, the impedanceparameter value, the equilibrium angle parameter value, a joint angle,and a joint angular velocity.
 12. The lower limb device of claim 1,wherein the knee and ankle joints are each configured in the propulsionphase to behave as a combined stiffness and damping, wherein thestiffness and damping components are configured such that the poweredknee and ankles joints synchronously generate energy associated withaccelerating the vertical motion of the body center of mass.